Methods for scanning an examination object using a CT system are generally known. These use for example circular scans, sequential circular scans with advance or spiral scans. During such scans at least one x-ray source and at least one opposing detector are used to record absorption data of the examination object from different recording angles and the absorption data or projections thus collected is/are processed by means of corresponding reconstruction methods to provide sectional images through the examination object.
To reconstruct computed tomography images from x-ray CT data records of a computed tomography device (CT device), i.e. from the captured projections, the standard method currently used is what is known as a filtered back projection or FBP method. Once the data has been captured, what is known as a rebinning step is carried out, in which the data generated with the beam propagated in the manner of a fan from the source is restructured so that it is present as if the detector were struck by x-ray beams arriving at the detector in a parallel manner. The data is then transformed to the frequency domain. Filtering takes place in the frequency domain and the filtered data is then back transformed. The data that has been reorganized and filtered thus is then used for a back projection onto the individual voxels within the volume of interest. However the approximate mode of operation of the conventional FBP methods results in problems with what are known as low-frequency cone beam artifacts and spiral artifacts.
In recent times therefore iterative reconstruction methods have been developed, which can eliminate at least some of the limitations of the FBP methods. With such an iterative reconstruction method initial image data from the projection measurement data is first reconstructed. A convolution back projection method for example can be used to this end. From this initial image data a “projector” or projection operator, which is intended to map the measurement system mathematically as effectively as possible, is used to generate synthetic projection data. The difference in relation to the measurement signals is then back projected using the operator adjoined to the projector and a residual image is thus reconstructed, which is used to update the initial image. The updated image data can in turn be used in a subsequent iteration step with the aid of the projection operator to generate new synthetic projection data, form the difference in relation to the measurement signals from this again and calculate a new residual image, which is again used to improve the image data of the current iteration stage, etc. Such a method allows image data to be reconstructed that has relatively good image sharpness but still has a low level of image noise.
One disadvantage of all calculation methods is that, when the examination object is moving or at least partially moving, motion blur can occur in the image, since during the time of a scanning operation for the data required for an image, a locational offset of the examination object or a part of the examination object may be present, so the measurement data producing an image does not all reflect a spatially identical situation of the examination object. This motion blur problem is particularly prevalent when carrying out cardio CT examinations of a patient, in which the movement of the heart can produce a high level of motion blur in the region of the heart or for examinations in which relatively fast changes in the examination object are to be measured.
The unwanted motion artifacts are reduced by increasing the temporal resolution of the CT recording. There are various procedures for this. On the one hand it is possible to reduce the gantry rotation time. However this soon encounters mechanical limitations, since the centrifugal force acting on the components increases quadratically as the rotation time decreases. On the other hand it is possible in the context of image reconstruction to improve temporal resolution by using in phase, complementary angle data of adjacent cardiac cycles. However the gain depends on the ratio of heart rate to gantry rotation time and cannot easily be influenced. Finally dual emitter CT systems have been developed, in other words CT devices with two x-ray sources and detectors assigned to these. The fact that measurement time is halved due to the presence of two x-ray source/detector systems allows double the temporal resolution. The disadvantage here is that the costs of the duplicated design of the core components, such as the emitter, detector, etc., are considerable.